Magnetic resonance imaging compatibility alloy for implantable medical devices

ABSTRACT

A biocompatible solid-solution alloy may be formed into any number of implantable medical devices. The solid-solution alloy comprises a combination of elements in specific ratios that make it magnetic resonance imaging compatible while retaining the characteristics required for implantable medical devices. The biocompatible solid-solution alloy is a cobalt-chromium alloy having substantially reduced iron, silicon, phosphorus and sulfur content.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to alloys for use in manufacturing orfabricating implantable medical devices, and more particularly, toimplantable medical devices manufactured or fabricated from alloys thatare magnetic resonance imaging compatible.

2. Discussion of the Related Art

Percutaneous transluminal angioplasty (PTA) is a therapeutic medicalprocedure used to increase blood flow through an artery. In thisprocedure, the angioplasty balloon is inflated within the stenosedvessel, or body passageway, in order to shear and disrupt the wallcomponents of the vessel to obtain an enlarged lumen. With respect toarterial stenosed lesions, the relatively incompressible plaque remainsunaltered, while the more elastic medial and adventitial layers of thebody passageway stretch around the plaque. This process producesdissection, or a splitting and tearing, of the body passageway walllayers, wherein the intima, or internal surface of the artery or bodypassageway, suffers fissuring. This dissection forms a “flap” ofunderlying tissue, which may reduce the blood flow through the lumen, orcompletely block the lumen. Typically, the distending intraluminalpressure within the body passageway can hold the disrupted layer, orflap, in place. If the intimal flap created by the balloon dilationprocedure is not maintained in place against the expanded intima, theintimal flap can fold down into the lumen and close off the lumen, ormay even become detached and enter the body passageway. When the intimalflap closes off the body passageway, immediate surgery is necessary tocorrect the problem.

Recently, transluminal prostheses have been widely used in the medicalarts for implantation in blood vessels, biliary ducts, ureters, or othersimilar organs of the living body. These prostheses are commonlyreferred to as stents and are used to maintain, open, or dilate tubularstructures. An example of a commonly used stent is given in U.S. Pat.No. 4,733,665 to Palmaz. Such stents are often referred to as balloonexpandable stents. Typically the stent is made from a solid tube ofstainless steel. Thereafter, a series of cuts are made in the wall ofthe stent. The stent has a first smaller diameter, which permits thestent to be delivered through the human vasculature by being crimpedonto a balloon catheter. The stent also has a second, expanded diameter,upon application of a radially, outwardly directed force, by the ballooncatheter, from the interior of the tubular shaped member.

However, one concern with such stents is that they are often impracticalfor use in some vessels such as the carotid artery. The carotid arteryis easily accessible from the exterior of the human body, and is closeto the surface of the skin. A patient having a balloon expandable stentmade from stainless steel or the like, placed in their carotid artery,might be susceptible to severe injury through day-to-day activity. Asufficient force placed on the patient's neck could cause the stent tocollapse, resulting in injury to the patient. In order to prevent this,self-expanding stents have been proposed for use in such vessels.Self-expanding stents act like springs and will recover to theirexpanded or implanted configuration after being crushed.

The prior art makes reference to the use of alloys such as Nitinol(Ni—Ti alloy), which have shape memory and/or superelasticcharacteristics, in medical devices, which are designed to be insertedinto a patient's body, for example, self-expanding stents. The shapememory characteristics allow the devices to be deformed to facilitatetheir insertion into a body lumen or cavity and then be heated withinthe body so that the device returns to its original shape. Superelasticcharacteristics, on the other hand, generally allow the metal to bedeformed and restrained in the deformed condition to facilitate theinsertion of the medical device containing the metal into a patient'sbody, with such deformation causing the phase transformation. Oncewithin the body lumen, the restraint on the superelastic member can beremoved, thereby reducing the stress therein so that the superelasticmember can return to its original un-deformed shape by thetransformation back to the original phase.

One concern with self-expanding stents and with other medical devicesformed from superelastic materials, is that they may exhibit reducedradiopacity under X-ray fluoroscopy. To overcome this problem, it iscommon practice to attach markers, made from highly radiopaquematerials, to the stent, or to use radiopaque materials in plating orcoating processes. Those materials typically include gold, platinum, ortantalum. The prior art makes reference to these markers or processes inU.S. Pat. No. 5,632,771 to Boatman et al., U.S. Pat. No. 6,022,374 toImran, U.S. Pat. No. 5,741,327 to Frantzen, U.S. Pat. No. 5,725,572 toLam et al., and U.S. Pat. No. 5,800,526 to Anderson et al. However, dueto the size of the markers and the relative position of the materialsforming the markers in the galvanic series versus the position of thebase metal of the stent in the galvanic series, there is a certainchallenge to overcome; namely, that of galvanic corrosion. Also, thesize of the markers increases the overall profile of the stent. Inaddition, typical markers are not integral to the stent and thus mayinterfere with the overall performance of the stent as well as becomedislodged from the stent.

A concern with both balloon expandable and self-expandable stents ismagnetic resonance imaging compatibility. Currently available metallicstents are known to cause artifacts in magnetic resonance generatedimages. In general, metals having a high magnetic permeability causeartifacts, while metals having a low magnetic permeability cause less orsubstantially no artifacts. In other words, if the stent or othermedical device is fabricated from a metal or metals having a lowmagnetic permeability, then less artifacts are created during magneticresonance imaging, which in turn allows more tissue in proximity to thestent or other medical device to be imaged.

Artifacts created under magnetic resonance imaging are promoted by localmagnetic field inhomogeneities and eddy currents induced by the magneticfield generated by the magnetic resonance imaging machine. The strengthof the magnetic field disruption is proportional to the magneticpermeability of the metallic stent or other medical device. In addition,signal attenuation within the stent is caused by radio frequencyshielding of the metallic stent or other medical device material.Essentially, the radio frequency signals generated by the magneticresonance imaging machine may become trapped within the cage likestructure of the stent or other medical device. Induced eddy currents inthe stent may also lead to a lower nominal radio frequency excitationangle inside the stent. This has been shown to attenuate the signalacquired by the receiver coil of the magnetic resonance imaging device.Artifact related signal changes may include signal voids or local signalenhancements, which in turn degrades the diagnostic value of the tool.

Accordingly, there is a need to develop materials for implantablemedical devices, such as stents, that are magnetic resonance imagingcompatible while retaining the toughness, durability and ductilityproperties required of implantable medical devices such as stents.

SUMMARY OF THE INVENTION

The present invention overcomes the diagnostic tool limitationsassociated with currently available implantable medical devices asbriefly described above.

In accordance with one aspect, the present invention is directed to animplantable medical device being formed from an improved, magneticresonance compatible solid-solution alloy. The solid solution alloycomprises chromium in the range from about 19 weight percent to about 21weight percent, tungsten in the range from about 14 weight percent toabout 16 weight percent, nickel in the range from about 9 weight percentto about 11 weight percent, manganese in the range from about 1 weightpercent to about 2 weight percent, carbon in the range from about 0.05weight percent to about 0.15 weight percent, iron in an amount not toexceed 0.3 weight percent, silicon in an amount not to exceed 0.4 weightpercent, phosphorus in an amount not to exceed 0.04 weight percent,sulfur in an amount not to exceed 0.03 weight percent and the remaindercobalt.

In accordance with another aspect, the present invention is directed toa biocompatible, load-carrying metallic structure being formed from animproved, magnetic resonance compatible solid solution alloy. The solidsolution alloy comprises chromium in the range from about 19 weightpercent to about 21 weight percent, tungsten in the range from about 14weight percent to about 16 weight percent, nickel in the range fromabout 9 weight percent to about 11 weight percent, manganese in therange from about 1 weight percent to about 2 weight percent, carbon inthe range from about 0.05 weight percent to about 0.15 weight percent,iron in an amount not to exceed 0.3 weight percent, silicon in an amountnot to exceed 0.4 weight percent, phosphorus in an amount not to exceed0.04 weight percent, sulfur in an amount not to exceed 0.03 weightpercent and the remainder cobalt.

The biocompatible alloy for implantable medical devices of the presentinvention offers a number of advantages over currently utilized alloys.The biocompatible alloy of the present invention is magnetic resonanceimaging compatible, is less brittle than other alloys, has enhancedductility and toughness, and has increased durability. The biocompatiblealloy also maintains the desired or beneficial characteristics ofcurrently available alloys including strength and flexibility.

The magnetic resonance imaging compatibility of implantable medicaldevices is gaining interest for the guidance of endovascularinterventional procedures and post-treatment evaluation. The magneticresonance imaging compatibility of the material or materials forming themedical devices is related to the basic magnetic susceptibility of thematerials relative to human tissue. A number of elements areferromagnetic, including iron, cobalt and nickel; however, iron has amagnetic susceptibility multiple orders of magnitude greater than theseother elements. Accordingly, reducing the iron content in an alloysubstantially reduces the magnetic susceptibility of the alloy, therebyenhancing magnetic resonance imaging.

The biocompatible alloy for implantable medical devices of the presentinvention may be utilized for any number of medical applications,including vessel patency devices such as vascular stents, biliarystents, ureter stents, vessel occlusion devices such as atrial septaland ventricular septal occluders, patent foramen ovale occluders andorthopedic devices such as fixation devices. In addition, thebiocompatible alloy may be utilized in the construction of deliverydevices for various medical devices. For example, the alloy may beutilized in the fabrication of guidewires.

The biocompatible alloy of the present invention is simple andinexpensive to manufacture. The biocompatible alloy may be formed intoany number of structures or devices. The biocompatible alloy may bethermomechanically processed, including cold-working and heat treating,to achieve varying degrees of strength and ductility. The biocompatiblealloy of the present invention may be age hardened to precipitate one ormore secondary phases.

BRIEF DESCRIPTION OF THE DRAWINGS

The foregoing and other features and advantages of the invention will beapparent from the following, more particular description of preferredembodiments of the invention, as illustrated in the accompanyingdrawings.

FIG. 1 is a graphical representation of the transition of criticalmechanical properties as a function of thermomechanical processing forcobalt-chromium alloys in accordance with the present invention.

FIG. 2 is a graphical representation of the endurance limit chart as afunction of thermomechanical processing for a cobalt-chromium alloy inaccordance with the present invention.

FIG. 3 is a flat layout diagrammatic representation of an exemplarystent fabricated from the biocompatible alloy in accordance with thepresent invention.

FIG. 4 is an enlarged view of the “M” links of the exemplary stent ofFIG. 3 in accordance with the present invention.

FIG. 5 is an enlarged view of a portion of the exemplary stent of FIG. 3in accordance with the present invention.

FIGS. 6 a and 6 b are magnetic resonance images of wires comprisingdifferent alloys in a magnetic field in accordance with the presentinvention.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

Biocompatible, solid-solution alloys may be utilized in the manufactureof any number of implantable medical devices. The biocompatible alloyfor implantable medical devices in accordance with the present inventionoffers a number of advantages over currently utilized medical gradealloys. In particular, the biocompatible alloy of the present inventionis magnetic resonance imaging compatible. Magnetic resonance imaging isa valuable diagnostic tool and thus any implantable medical deviceshould preferably be magnetic resonance imaging compatible so that itand surrounding tissue may be accurately imaged. One such medical devicewhere this is particularly relevant is stents.

Coronary stenting is currently the most widely utilized percutaneouscoronary intervention. The primary benefit of stenting when comparedwith balloon-angioplasty alone is a reduction of the restenosis-rate.Nevertheless, in-stent restenosis remains a relatively common clinicalscenario. If clinical symptoms suggest in-stent restenosis, x-raycoronary angiography is currently considered the standard for theevaluation of stent integrity. Conventional x-ray angiography has anumber of disadvantages, including a small risk of potentially seriouscomplications, the need for a contrast agent containing a form ofiodine, and radiation exposure. Accordingly, a noninvasive imagingmethod for direct assessment of stent lumen integrity would bepreferable. Magnetic resonance imaging provides such a method.

Currently, the majority of coronary artery stents are ferromagnetic butare considered to be magnetic resonance imaging safe. Although thesedevices are considered magnetic resonance imaging safe, theytraditionally induce image artifacts that may pose inaccurate,clinically relevant inferences when inspected by a clinician. Forexample, traditional biocompatible cobalt-alloys such as L605 (commontradename: Haynes 25 from the Haynes International Corporation) can haveas much as 3 wt. % iron. Biocompatible metallic alloys that containstrongly ferromagnetic materials such as iron, but not limited thereto,generally exhibit a high magnetic permeability which tends to induceunintended image artifacts. Moreover, traditional biocompatibleferrous-based alloys such as stainless steel may contain significantlygreater concentrations of strongly ferromagnetic materials such as iron.

For reference, a traditional stainless steel alloy such as 316L (i.e.UNS S31603) which is broadly utilized as an implantable, biocompatibledevice material may comprise chromium (Cr) in the range from about 16 to18 wt. %, nickel (Ni) in the range from about 10 to 14 wt. %, molybdenum(Mo) in the range from about 2 to 3 wt. %, manganese (Mn) in the rangeup to 2 wt. %, silicon (Si) in the range up to 1 wt. %, with iron (Fe)comprising the balance (approximately 65 wt. %) of the composition.

Additionally, a traditional cobalt-based alloy such as L605 (i.e. UNSR30605) which is also broadly utilized as an implantable, biocompatibledevice material may comprise chromium (Cr) in the range from about 19 to21 wt. %, tungsten (W) in the range from about 14 to 16 wt. %, nickel(Ni) in the range from about 9 to 11 wt. %, iron (Fe) in the range up to3 wt. %, manganese (Mn) in the range up to 2 wt. %, silicon (Si) in therange up to 1 wt. %, with cobalt (Co) comprising the balance(approximately 49 wt. %) of the composition.

In general, elemental additions such as chromium (Cr), nickel (Ni),tungsten (W), manganese (Mn), silicon (Si) and molybdenum (Mo) whereadded to iron- and/or cobalt-based alloys, where appropriate, toincrease or enable desirable performance attributes, including strength,machinability and corrosion resistance within clinically relevant usageconditions.

The composition of the material of the present invention does noteliminate ferromagnetic components but rather shift the ‘susceptibility’(i.e. the magnetic permeability) such that the magnetic resonanceimaging compatibility may be improved. In addition, the material of thepresent invention is intended to improve the measurable ductility byminimizing the deleterious effects induced by traditional machiningaides such as silicon (Si).

The traditional cobalt-based alloy, L605, is a nonmagneticchromium-nickel-tungsten-cobalt alloy. Among the elements comprising theL605 alloy, iron, cobalt and nickel are known ferromagnetic metals. Ofthese three elements, iron has the highest magnetic susceptibilitylevel. Magnetic susceptibility is a unitless constant that is determinedby the physical properties of the material. More particularly, iron hasa magnetic susceptibility level of 200,000 in c.g.s. units, cobalt has amagnetic susceptibility level of 250 in c.g.s. units, and nickel has amagnetic susceptibility level of 600 in c.g.s. units. These magneticsusceptibility levels indicate that the iron in the L605 alloy may bethe most influential element to the overall L605 alloy's magneticproperties. While the detailed magnetic susceptibility of L605 isunclear, the magnetic rating for L605 is estimated to be within therange between paramagnetic to ferromagnetic.

The iron content in the L605 alloy is a maximum of 3 weight percent. Inaccordance with an exemplary embodiment, the iron content of the alloymay be reduced to a level of 1 percent or less, and more particularly toa level of less than 0.3 weight percent, with the reduction beingcovered by an increase in cobalt. The variation in weight percent ofiron and cobalt as set forth herein, does not have a measurable impacton material mechanical properties. Accordingly, by controlling themanufacturing process as described herein, it is possible andeconomically practical to produce an alloy with a significantly reducediron content, thereby reducing the overall magnetic susceptibility ofthe alloy.

Referring now to FIG. 6 a, there is illustrated a magnetic resonanceimage of two wires positioned such that they are substantiallyperpendicular to the magnetic field. Wire 602 is a 0.005 inch diameterwire formed from standard L605 (iron content of approximately 2.29percent by weight). Wire 604 is a 0.005 inch diameter wire formed fromthe improved alloy (low iron of approximately 0.1 percent by weight) ofthe present invention. As may be seen from a comparison of the twoimages, the low iron alloy of the present invention results in a clearerimage. In FIG. 6 b, the same wires are positioned substantially parallelto the magnetic field. Wire 606, which is formed from standard L605, isblurry, while wire 608 appears substantially invisible. A conclusionbased on the two sets of images may be that the low iron alloy of thepresent invention induces less artifacts and results in better images.

In accordance with an exemplary embodiment, an implantable medicaldevice may be formed from a solid-solution alloy comprising chromium inthe range from about 19 weight percent to about 21 weight percent,tungsten in the range from about 14 weight percent to about 16 weightpercent, nickel in the range from about 9 weight percent to about 11weight percent, manganese in the range from about 1 weight percent toabout 2 weight percent, carbon in the range from about 0.05 weightpercent to about 0.15 weight percent, iron in an amount not to exceed0.3 weight percent, silicon in an amount not to exceed 0.4 weightpercent, phosphorus in an amount not to exceed 0.04 weight percent,sulfur in an amount not to exceed 0.03 weight percent and the remaindercobalt.

In contrast to the traditional formulation of this alloy (i.e.L605/Haynes 25), the intended composition does not include any elementaliron (Fe) or silicon (Si) above conventional accepted trace impuritylevels. Accordingly, this exemplary embodiment will exhibit a markedreduction in ‘susceptibility’ (i.e. the magnetic permeability) therebyleading to improved magnetic resonance imaging compatibility.Additionally, the exemplary embodiment will exhibit a marked improvementin material ductility and fatigue strength (i.e. cyclic endurance limitstrength) due to the elimination of silicon (Si), above trace impuritylevels.

The preferred embodiment may be processed from the requisite elementaryraw materials, as set-forth above, by first mechanical homogenization(i.e. mixing) and then compaction into a green state (i.e. precursory)form. If necessary, appropriate manufacturing aids such as hydrocarbonbased lubricants and/or solvents (e.g. mineral oil, machine oils,kerosene, isopropanol and related alcohols) be used to ensure completemechanical homogenization. Additionally, other processing steps such asultrasonic agitation of the mixture followed by cold compaction toremove any unnecessary manufacturing aides and to reduce void spacewithin the green state may be utilized. It is preferable to ensure thatany impurities within or upon the processing equipment from priorprocessing and/or system construction (e.g. mixing vessel material,transfer containers, etc.) be sufficiently reduced in order to ensurethat the green state form is not unnecessarily contaminated. This may beaccomplished by adequate cleaning of the mixing vessel before adding theconstituent elements by use of surfactant based cleaners to remove anyloosely adherent contaminants.

Initial melting of the green state form into a ingot of desiredcomposition, is achieved by vacuum induction melting (VIM) where theinitial form is inductively heated to above the melting point of theprimary constituent elements within a refractory crucible and thenpoured into a secondary mold within a vacuum environment (e.g. typicallyless than or equal to 10⁻⁴ mmHg). The vacuum process ensures thatatmospheric contamination is significantly minimized. Uponsolidification of the molten pool, the ingot bar is substantially singlephase (i.e. compositionally homogenous) with a definable threshold ofsecondary phase impurities that are typically ceramic (e.g. carbide,oxide or nitride) in nature. These impurities are typically inheritedfrom the precursor elemental raw materials.

A secondary melting process termed vacuum arc reduction (VAR) isutilized to further reduce the concentration of the secondary phaseimpurities to a conventionally accepted trace impurity level (i.e.<1,500 ppm). Other methods maybe enabled by those skilled in the art ofingot formulation that substantially embodies this practice of ensuringthat atmospheric contamination is minimized. In addition, the initialVAR step may be following followed by repetitive VAR processing tofurther homogenize the solid-solution alloy in the ingot form. From theinitial ingot configuration, the homogenized alloy will be furtherreduced in product size and form by various industrially acceptedmethods such as, but not limited too, ingot peeling, grinding, cutting,forging, forming, hot rolling and/or cold finishing processing steps soas to produce bar stock that may be further reduced into a desired rawmaterial form.

In this exemplary embodiment, the initial raw material product form thatis required to initiate the thermomechanical processing that willultimately yield a desired small diameter, thin-walled tube, appropriatefor interventional devices, is a modestly sized round bar (e.g. one inchin diameter round bar stock) of predetermined length. In order tofacilitate the reduction of the initial bar stock into a much smallertubing configuration, an initial clearance hole must be placed into thebar stock that runs the length of the product. These tube hollows (i.e.heavy walled tubes) may be created by ‘gun-drilling’ (i.e. high depth todiameter ratio drilling) the bar stock. Other industrially relevantmethods of creating the tube hollows from round bar stock may beutilized by those skilled-in-the-art of tube making.

Consecutive mechanical cold-finishing operations such as drawing througha compressive outer-diameter (OD), precision shaped (i.e. cut),circumferentially complete, diamond die using any of the followinginternally supported (i.e. inner diameter, ID) methods, but notnecessarily limited to these conventional forming methods, such as hardmandrel (i.e. relatively long traveling ID mandrel also referred to asrod draw), floating-plug (i.e. relatively short ID mandrel that ‘floats’within the region of the OD compressive die and fixed-plug (i.e. the IDmandrel is ‘fixed’ to the drawing apparatus where relatively shortworkpieces are processed) drawing. These process steps are intended toreduce the outer-diameter (OD) and the corresponding wall thickness ofthe initial tube hollow to the desired dimensions of the finishedproduct.

When necessary, tube sinking (i.e. OD reduction of the workpiece withoutinducing substantial tube wall reduction) is accomplished by drawing theworkpiece through a compressive die without internal support (i.e. no IDmandrel). Conventionally, tube sinking is typically utilized as a finalor near-final mechanical processing step to achieve the desireddimensional attributed of the finished product.

Although not practically significant, if the particular compositionalformulation will support a single reduction from the initial rawmaterial configuration to the desired dimensions of the finishedproduct, in process heat-treatments will not be necessary. Wherenecessary to achieve intended mechanical properties of the finishedproduct, a final heat-treating step is utilized.

Conventionally, all metallic alloys in accordance with the presentinvention will require incremental dimensional reductions from theinitial raw material configuration to reach the desired dimensions ofthe finished product. This processing constraint is due to thematerial's ability to support a finite degree of induced mechanicaldamage per processing step without structural failure (e.g.strain-induced fracture, fissures, extensive void formation, etc.).

In order to compensate for induced mechanical damage (i.e. cold-working)during any of the aforementioned cold-finishing steps, periodic thermalheat-treatments are utilized to stress-relieve (i.e. minimization ofdeleterious internal residual stresses that are the result of processessuch as cold-working) thereby increasing the workability (i.e. abilityto support additional mechanical damage without measurable failure) theworkpiece prior to subsequent reductions. These thermal treatments aretypically, but not necessarily limited to, conducted within a relativelyinert environment such as an inert gas furnace (e.g. nitrogen, argon,etc.), a oxygen rarified hydrogen furnace, a conventional vacuum furnaceand under less common process conditions, atmospheric air. When vacuumfurnaces are utilized, the level of vacuum (i.e. subatmosphericpressure), typically measured in units of mmHg or torr (where 1 mmHg isequal to 1 unit torr), shall be sufficient to ensure that excessive anddeteriorative high temperature oxidative processes are not functionallyoperative during heat treatment. This process may usually be achievedunder vacuum conditions of 10⁻⁴ mmHg (0.0001 torr) or less (i.e. lowermagnitude).

The stress relieving heat treatment temperature is typically heldconstant between 82 to 86% of the conventional melting point (i.e.industrially accepted liquidus temperature, 0.82 to 0.86 homologoustemperatures) within an adequately sized isothermal region of theheat-treating apparatus. The workpiece undergoing thermal treatment isheld within the isothermal processing region for a finite period of timethat is adequate to ensure that the workpiece has reached a state ofthermal equilibrium and for that sufficient time is elapsed to ensurethat the reaction kinetics (i.e. time dependent material processes) ofstress-relieving and/or process annealing, as appropriate, is adequatelycompleted. The finite amount of time that the workpiece is held withinthe processing is dependent upon the method of bringing the workpieceinto the process chamber and then removing the working upon completionof heat treatment. Typically, this process is accomplished by, but notlimited to, use of a conventional conveyor-belt apparatus or otherrelevant mechanical assist devices. In the case of the former, theconveyor belt speed and appropriate finite dwell-time, as necessary,within the isothermal region is controlled to ensure that sufficienttime at temperature is utilized so as to ensure that the process iscompleted as intended.

When necessary to achieve desired mechanical attributes of the finishedproduct, heat-treatment temperatures and corresponding finite processingtimes may be intentionally utilized that are not within the typicalrange of 0.82 to 0.86 homologous temperatures. Various age hardening(i.e. a process that induces a change in properties at moderatelyelevated temperatures, relative to the conventional melting point, thatdoes not induce a change in overall chemical composition change in themetallic alloy being processed) processing steps may be carried out, asnecessary, in a manner consistent with those previously described attemperatures substantially below 0.82 to 0.86 homologous temperature.For Co-based alloys in accordance with the present invention, theseprocessing temperatures may be varied between and inclusive ofapproximately 0.29 homologous temperature and the aforementioned stressrelieving temperature range. The workpiece undergoing thermal treatmentis held within the isothermal processing region for a finite period oftime that is adequate to ensure that the workpiece has reached a stateof thermal equilibrium and for that sufficient time is elapsed to ensurethat the reaction kinetics (i.e. time dependent material processes) ofage hardening, as appropriate, is adequately completed prior to removalfrom the processing equipment.

In some cases for Co-based alloys in accordance with the presentinvention, the formation of secondary-phase ceramic compounds such ascarbide, nitride and/or oxides will be induced or promoted by agehardening heat treating. These secondary-phase compounds are typically,but not limited to, for Co-based alloys in accordance with the presentinvention, carbides which precipitate along thermodynamically favorableregions of the structural crystallographic planes that comprise eachgrain (i.e. crystallographic entity) that make-up the entirepolycrystalline alloy. These secondary-phase carbides can exist alongthe intergranular boundaries as well as within each granular structure(i.e. intragranular). Under most circumstances for Co-based alloys inaccordance with the present invention, the principal secondary phasecarbides that are stoichiometrically expected to be present are M₆Cwhere M typically iscobalt (Co). When present, the intermetallic M₆Cphase is typically expected to reside intragranularly alongthermodynamically favorable regions of the structural crystallographicplanes that comprise each grain within the polycrystalline alloy inaccordance with the present invention. Although not practically common,the equivalent material phenomena can exist for a single crystal (i.e.monogranular) alloy.

Additionally, another prominent secondary phase carbide can also beinduced or promoted as a result of age hardening heat treatments. Thisphase, when present, is stoichiometrically expected to be M₂₃C₆ where Mtypically is chromium (Cr) but is also commonly observed to be cobalt(Co) especially in Co-based alloys. When present, the intermetallicM₂₃C₆ phase is typically expected to reside along the intergranularboundaries (i.e. grain boundaries) within a polycrystalline alloy inaccordance with the present invention. As previously discussed for theintermetallic M₆C phase, the equivalent presence of the intermetallicM₂₃C₆ phase can exist for a single crystal (i.e. monogranular) alloy,albeit not practically common.

In the case of the intergranular M₂₃C₆ phase, this secondary phase isconventionally considered most important, when formed in a manner thatis structurally and compositionally compatible with the alloy matrix, tostrengthening the grain boundaries to such a degree that intrinsicstrength of the grain boundaries and the matrix are adequately balanced.By inducing this equilibrium level of material strength at themicrostructural level, the overall mechanical properties of the finishedtubular product can be further optimized to desirable levels.

In addition to stress relieving and age hardening related heat-treatingsteps, solutionizing (i.e. sufficiently high temperature and longerprocessing time to thermodynamically force one of more alloyconstituents to enter into solid solution—‘singular phase’, alsoreferred to as full annealing) of the workpiece may be utilized. ForCo-based alloys in accordance with the present invention, the typicalsolutionizing temperature can be varied between and inclusive ofapproximately 0.88 to 0.90 homologous temperatures. The workpieceundergoing thermal treatment is held within the isothermal processingregion for a finite period of time that is adequate to ensure that theworkpiece has reached a state of thermal equilibrium and for thatsufficient time is elapsed to ensure that the reaction kinetics (i.e.time dependent material processes) of solutionizing, as appropriate, isadequately completed prior to removal from the processing equipment.

The sequential and selectively ordered combination of thermomechanicalprocessing steps that may comprise but not necessarily includemechanical cold-finishing operations, stress relieving, age hardeningand solutionizing can induce and enable a broad range of measurablemechanical properties as a result of distinct and determinablemicrostructural attributes. This material phenomena can be observed inFIG. 1. which shows a chart that exhibits the affect of thermomechanicalprocessing (TMP) such as cold working and in-process heat-treatments onmeasurable mechanical properties such as yield strength and ductility(presented in units of percent elongation) in accordance with thepresent invention. In this example, thermomechanical (TMP) groups one(1) through five (5) were subjected to varying combinations ofcold-finishing, stress relieving and age hardening and not necessarilyin the presented sequential order. In general, the principal isothermalage hardening heat treatment applied to each TMP group varied betweenabout 0.74 to 0.78 homologous temperatures for group (1), about 0.76 to0.80 homologous temperatures for group (2), about 0.78 to 0.82homologous temperatures for group (3), about 0.80 to 0.84 homologoustemperatures for group (4) and about 0.82 to 0.84 homologoustemperatures for group (5). The each workpiece undergoing thermaltreatment was held within the isothermal processing region for a finiteperiod of time that was adequate to ensure that the workpiece reached astate of thermal equilibrium and to ensure that sufficient time waselapsed to ensure that the reaction kinetics of age hardening wasadequately completed.

More so, the effect of thermomechanical (TMP) on cyclic fatigueproperties is on Co-based alloys, in accordance with the presentinvention, is reflected in FIG. 2. Examination of FIG. 2. shows theaffect on fatigue strength (i.e. endurance limit) as a function ofthermomechanical processing for the previously discussed TMP groups (2)and (4). TMP group (2) from this figure as utilized in this specificexample shows a marked increase in the fatigue strength (i.e. endurancelimit, the maximum stress below which a material can presumably endurean infinite number of stress cycles) over and against the TMP group (4)process.

The above-described alloy may be utilized in any number of implantablemedical devices. The alloy is particularly advantageous in situationswhere magnetic resonance imaging is a useful diagnostic tool such asdetermining in-stent restenosis. Accordingly, although the alloy may beutilized for any implantable medical device, an exemplary stentconstructed from the alloy is described below.

FIG. 3 is a flat layout of an exemplary embodiment of a stent that maybe constructed utilizing the alloy of the present invention. The stent10 comprises end sets of strut members 12 located at each end of thestent 10 and central sets of strut members 14 connected each to theother by sets of flexile “M” links 16. Each end set of strut members 12comprises alternating curved sections 18 and diagonal sections 20connected together to form a closed circumferential structure. Thecentral sets of strut members 14 located longitudinally between the endsets of strut members 14 comprise curved sections 22 and diagonalsections 24 connected together to form a closed circumferentialring-like structure.

Referring to FIG. 4 there is illustrated an enlargement of the flexible“M” links 16 of the stent 10. Each “M” link 16 has a circumferentialextent, i.e. length, L′ above and L″ below line 11. The line 11 is drawnbetween the attachment points 13 where the “M” link 16 attaches toadjacent cured sections 18 or 22. Such a balanced design preferablydiminishes any likelihood of the flexible connecting link 16 fromexpanding into the lumen of artery or other vessel.

As illustrated in FIG. 3, the diagonal sections 20 of the end sets ofstrut members 12 are shorter in length than the diagonal sections 24 ofthe central sets of strut members 14. The shorter diagonal sections 20will preferably reduce the longitudinal length of metal at the end ofthe stent 10 to improve deliverability into a vessel of the human body.In the stent 10, the widths of the diagonal sections 20 and 24 aredifferent from one another.

Referring to FIG. 5, there is illustrated an expanded view of a stentsection comprising an end set of strut members 12 and a central set ofstrut members 14. As illustrated, the diagonal sections 24 of thecentral sets of strut members 14 have a width at the center thereof,T_(c), and a width at the end thereof, T_(e), wherein T_(c) is greaterthan T_(e). This configuration allows for increased radiopacity withoutaffecting the design of curved sections 22 that are the primary stentelements involved for stent expansion. In an exemplary embodiment, thecurved sections 22 and 18 may be tapered and may have uniform widthswith respect to one another as is explained in detail subsequently. Thediagonal sections 20 of the end sets of strut members 12 also have atapered shape. The diagonal sections 20 have a width in the center,T_(c)-end, and a width at the end, T_(e)-end, wherein T_(c)-end isgreater than T_(e)-end. Because it is preferable for the end sets ofstrut members 12 to be the most radiopaque part of the stent 10, thediagonal section 20 center width T_(c)-end of the end sets of strutmembers 12 is wider than the width T_(c) of the diagonal section 24.Generally, a wider piece of metal will be more radiopaque. Thus, thestent 10 has curved sections with a single bend connecting the diagonalsections of its sets of strut members, and flexible connecting linksconnecting the curved sections of its circumferential sets of strutmembers.

The width of the curved sections 22 and 18 taper down as one moves awayfrom the center of the curve until a predetermined minimum widthsubstantially equal to that of their respective diagonal sections 24 and20. To achieve this taper, the inside arc of the curved sections 22 and18 have a center that is longitudinally displaced from the center of theoutside arc. This tapered shape for the curved sections 22 and 18provides a significant reduction in metal strain with little effect onthe radial strength of the expanded stent as compared to a stent havingsets of strut members with a uniform strut width.

This reduced strain design has several advantages. First, it can allowthe exemplary design to have a much greater usable range of radialexpansion as compared to a stent with a uniform strut width. Second, itcan allow the width at the center of the curve to be increased whichincreases radial strength without greatly increasing the metal strain(i.e. one can make a stronger stent). Finally, the taper reduces theamount of metal in the stent and that should improve the stentthrombogenicity.

The curved sections 18 of the end sets of strut members 12 and thecurved sections 22 of the central sets of strut members 14 have the samewidths. As a result of this design, the end sets of strut members 12,which have shorter diagonal sections 20, will reach the maximumallowable diameter at a level of strain that is greater than the levelof strain experienced by the central sets of strut members 14.

It is important to note that although a stent is described, the alloymay be utilized for any number of implantable medical devices.

Although shown and described is what is believed to be the mostpractical and preferred embodiments, it is apparent that departures fromspecific designs and methods described and shown will suggest themselvesto those skilled in the art and may be used without departing from thespirit and scope of the invention. The present invention is notrestricted to the particular constructions described and illustrated,but should be constructed to cohere with all modifications that may fallwithin the scope for the appended claims.

1. An implantable medical device being formed from an improved, magneticresonance compatible solid-solution alloy comprising chromium in therange from about 19 weight percent to about 21 weight percent, tungstenin the range from about 14 weight percent to about 16 weight percent,nickel in the range from about 9 weight percent to about 11 weightpercent, manganese in the range from about 1 weight percent to about 2weight percent, carbon in the range from about 0.05 weight percent toabout 0.15 weight percent, iron in an amount not to exceed 0.3 weightpercent, silicon in an amount not to exceed 0.4 weight percent,phosphorus in an amount not to exceed 0.04 weight percent, sulfur in anamount not to exceed 0.03 weight percent and the remainder cobalt. 2.The implantable medical device according to claim 1, wherein thesolid-solution alloy is constructed through thermomechanical processingto exhibit relatively high strength and low ductility characteristics inthe fully cold-worked state.
 3. The implantable medical device accordingto claim 1, wherein the solid-solution alloy is constructed throughthermomechanical processing to exhibit relatively moderate strength andmoderate ductility characteristics in the partially cold-worked state.4. The implantable medical device according to claim 3, wherein thesolid-solution alloy is further constructed through age hardening for apredetermined time within a gaseous environment at a temperature lessthan the annealing temperature to precipitate one or more secondaryphases, including at least one of intragranular and intergranularphases, from a substantially single phase structure.
 5. The implantablemedical device according to claim 4, wherein the age hardeningtemperature is in the range from about 750 degrees Fahrenheit to about2,150 degrees Fahrenheit.
 6. The implantable medical device according toclaim 4, wherein the age hardening gaseous environment compriseshydrogen, nitrogen, argon and air.
 7. The implantable medical deviceaccording to claim 3, wherein the solid-solution alloy is furtherconstructed through stress relieving for a predetermined time within agaseous environment at a temperature less than the annealing temperaturewhile maintaining a substantially single phase to increase toughness andductility.
 8. The implantable medical device according to claim 7,wherein the stress relieving temperature is about or less than 100degrees Fahrenheit below the annealing temperature.
 9. The implantablemedical device according to claim 7, wherein the stress relievinggaseous environment comprises hydrogen, nitrogen, argon and air.
 10. Theimplantable medical device according to claim 1, wherein thesolid-solution alloy is constructed through thermomechanical processingto exhibit relatively low strength and high ductility characteristics inthe fully annealed state.
 11. The implantable medical device accordingto claim 1, wherein the medical device comprises a stent.
 12. Theimplantable medical device according to claim 11, wherein the medicaldevice comprises a vascular stent.
 13. The implantable medical deviceaccording to claim 3, wherein the solid-solution alloy is furtherconstructed through stress relieving for a predetermined time with avacuum environment at a temperature less than the annealing temperaturewhile maintaining a substantially single phase to increase toughness andductility.
 14. The implantable medical device according to claim 13,wherein the stress relieving temperature is about or less than onehundred degrees Fahrenheit below the annealing temperature.
 15. Abiocompatible, load-carrying metallic structure being formed from animproved, magnetic resonance compatible solid-solution alloy comprisingchromium in the range from about 19 weight percent to about 21 weightpercent, tungsten in the range from about 14 weight percent to about 16weight percent, nickel in the range from about 9 weight percent to about11 weight percent, manganese in the range from about 1 weight percent toabout 2 weight percent, carbon in the range from about 0.05 weightpercent to about 0.15 weight percent, iron in an amount not to exceed0.3 weight percent, silicon in an amount not to exceed 0.4 weightpercent, phosphorus in an amount not to exceed 0.04 weight percent,sulfur in an amount not to exceed 0.03 weight percent and the remaindercobalt.
 16. The biocompatible, load-carrying metallic structureaccording to claim 15, wherein the solid-solution alloy is constructedthrough thermomechanical processing to exhibit relatively high strengthand low ductility characteristics in the fully cold-worked state. 17.The biocompatible, load-carrying metallic structure according to claim15, wherein the solid-solution alloy is constructed throughthermomechanical processing to exhibit relatively moderate strength andmoderate ductility characteristics in the partially cold-worked state.18. The biocompatible, load-carrying metallic structure according toclaim 15, wherein the solid-solution alloy is further constructedthrough age hardening for a predetermined time within a gaseousenvironment at a temperature less than the annealing temperature toprecipitate one or more secondary phases, including at least one ofintragranular and intergranular phases, from a substantially singlephase structure.
 19. The biocompatible, load-carrying metallic structureaccording to claim 18, wherein the age hardening temperature is in therange from about 750 degrees Fahrenheit to about 2,150 degreesFahrenheit.
 20. The biocompatible, load-carrying metallic structureaccording to claim 18, wherein the age hardening gaseous environmentcomprises hydrogen, nitrogen, argon and air.
 21. The biocompatible,load-carrying metallic structure according to claim 17, wherein thesolid-solution alloy is further constructed through stress relieving fora predetermined time within a gaseous environment at a temperature lessthan the annealing temperature while maintaining a substantially singlephase to increase toughness and ductility.
 22. The biocompatible,load-carrying metallic structure according to claim 21, wherein thestress relieving temperature is about or less than 100 degreesFahrenheit below the annealing temperature.
 23. The biocompatible,load-carrying metallic structure according to claim 21, wherein thestress relieving gaseous environment comprises hydrogen, nitrogen, argonand air.
 24. The biocompatible, load-carrying metallic structureaccording to claim 15, wherein the solid-solution alloy is constructedthrough thermomechanical processing to exhibit relatively low strengthand high ductility characteristics in the fully annealed state.
 25. Thebiocompatible, load-carrying metallic structure according to claim 15,wherein the medical device comprises a fixation device.
 26. Thebiocompatible, load-carrying metallic structure according to claim 15,wherein the medical device comprises an artificial joint implant. 27.The biocompatible, load-carrying metallic structure wherein thesolid-solution alloy is further constructed through stress relieving fora predetermined time with a vacuum environment at a temperature lessthan the annealing temperature while maintaining a substantially singlephase to increase toughness and ductility.
 28. The biocompatible,load-carrying metallic structure according to claim 27, wherein thestress relieving temperature is about or less than one hundred degreesFahrenheit below the annealing temperature.